The use of gamma ray detectors in general, and positron emission tomography (PET) in particular, is growing in the field of medical imaging. In PET imaging, a radiopharmaceutical agent is introduced into an object to be imaged via injection, inhalation, or ingestion. After administration of the radiopharmaceutical, the physical and bio-molecular properties of the agent will cause it to concentrate at specific locations in the human body. The actual spatial distribution of the agent, the intensity of the region of accumulation of the agent, and the kinetics of the process from administration to eventually elimination are all factors that may have clinical significance. During this process, a positron emitter attached to the radiopharmaceutical agent will emit positrons according to the physical properties of the isotope, such as half-life, branching ratio, etc.
The radionuclide emits positrons, and when an emitted positron collides with an electron, an annihilation event occurs, wherein the positron and electron are destroyed. Most of the time, an annihilation event produces two gamma rays (at 511 keV) traveling at substantially 180 degrees apart.
By detecting the two gamma rays, and drawing a line between their locations, i.e., the line-of-response (LOR), one can retrieve the likely location of the original disintegration. While this process will only identify a line of possible interaction, by accumulating a large number of those lines, and through a tomographic reconstruction process, the original distribution can be estimated. In addition to the location of the two scintillation events, if accurate timing (within few hundred picoseconds) is available, a time-of-flight (TOF) calculation can add more information regarding the likely position of the event along the line. Limitations in the timing resolution of the scanner will determine the accuracy of the positioning along this line. Limitations in the determination of the location of the original scintillation events will determine the ultimate spatial resolution of the scanner, while the specific characteristics of the isotope (e.g., energy of the positron) will also contribute (via positron range and co-linearity of the two gamma rays) to the determination of the spatial resolution the specific agent.
The above described detection process must be repeated for a large number of annihilation events. While each imaging case must be analyzed to determine how many counts (i.e., paired events) are required to support the imaging task, current practice dictates that a typical 100-cm long, FDG (fluoro-deoxyglucose) study will need to accumulate several hundred million counts. The time required to accumulate this number of counts is determined by the injected dose of the agent and the sensitivity and counting capacity of the scanner.
PET imaging systems use detectors positioned across from one another to detect the gamma rays emitting from the object. Typically a ring of detectors is used in order to detect gamma rays coming from each angle. Thus, a PET scanner is typically substantially cylindrical to be able to capture as much radiation as possible, which should be, by definition, isotropic. Once the overall geometry of the PET scanner is known, another challenge is to arrange as much scintillating material as possible in the gamma ray paths to stop and convert as many gamma rays as possible into light. In order to be able to reconstruct the spatio-temporal distribution of the radio-isotope via tomographic reconstruction principles, each detected event will need to be characterized for its energy (i.e., amount of light generated), its location, and its timing. Most modern PET scanners are composed of several thousand individual crystals, which are arranged in modules and are used to identify the position of the scintillation event. Typically crystal elements have a cross section of roughly 4 mm×4 mm. Smaller or larger dimensions and non-square sections are also possible. The length or depth of the crystal will determine how likely the gamma ray will be captured, and typically ranges from 10 to 30 mm. The detector module is the main building block of the scanner.
PET imaging relies on the conversion of gamma rays into light through fast and bright scintillation crystals. After determining the interaction position in the scintillator and time pairing of individual events, the location of the annihilation process can be recreated. These actions require very fast components—detector and electronics—and they also require excellent signal to noise ratio. With high quality electronics, the signal to noise ratio is mainly determined by the inherent Poisson statistics involved in the detection process. Detecting more photons will result in improved signal-to-noise-ratio, and, therefore, better spatial and timing resolution. No improvement in detector design and electronics can compensate for significant loss of light in the detection process. The fraction of the total amount of light collected (relative to the amount created in the scintillator) is a good measure of the efficiency of the design. So to maximize the amount of light collected, one would try to get the light sensor as close as possible to the scintillation crystal and avoid reflections and other edge effects. This would then force the arrangement to be large array detector with short distance between crystal and sensor.
As described above, a PET imaging system is more than just a counter and, in addition to detecting the presence of a scintillation event, the system must identify its location. Conceptually, perhaps the most straightforward design to allow identification of the location of each interaction is to have a separate photosensor and data acquisition channel for each scintillator crystal. Due to constraints such as the physical size of common photosensors, the power required for each data acquisition channel, and the associated cost of these items, some form of multiplexing is usually used to reduce the number of photosensors and channels of electronics.
By properly documenting how light is being distributed to the multiple light sensors, it is possible to assign an event location for any given set of sensor responses. Light therefore needs to be distributed to multiple sensors. In order to accomplish an adequate light distribution (so that enough sensors would detect a fraction of the light) it may be necessary to increase the thickness of the light guide or space between the crystals and the sensor. However, fast counting requires that multiple events be processed simultaneously, favoring optical isolation between scintillation events, and the creation of smaller detector blocks. These two requirements are pushing the detector design in two different directions.
Currently available PET scanners have two main detector module designs. The first type is a large area detector in which an array of crystals that covers the entire axial extent of the cylinder is formed. Several modules are then arranged together to form a cylinder, each module being optically coupled to the next. An array of photosensors (e.g., photomultiplier tubes or PMTs) is placed on the modules and on the interfaces between modules. See the design shown in FIG. 1A, which illustrates a module that includes an array of crystal elements and an array of PMTs. This approach minimizes the number of optical interfaces and boundaries, and ensures excellent light collection. However, this design suffers from larger numbers of sensors being exposed to the light of a single scintillation event, potentially limiting the ability to process events occurring close to each other, as well as limiting the overall counting capacity.
The second design is based upon an optically isolated block having, for example, four PMT sensors, so as to allow for simplified crystal identification. In the design of FIG. 1A, a block element is composed of four photomultiplier tube sensors on an approximately 50 mm×50 mm crystal assembly. In this approach, the crystals extend to the very edge of the array and a relatively thick light guide is therefore often used to capture enough light from all PMTs to be able to detect the position of the event. A detector is then formed by arranging multiple blocks (e.g., three or four) to fill out the axial extent, and then repeating this pattern to create the overall cylinder. See the designs shown in FIGS. 1B and 1C. The advantages of this approach include greater flexibility (the detector block is potentially fully functional outside of the scanner (meaning that the detector block can be tested and calibrated separately, as opposed to the large area, continuous detector that can only be tested and calibration as part of a complete system—offering advantages for service at the customer site and for manufacturing of the scanner)) and better count capacity due to the potential parallel operation of each module. The disadvantages of this design are the inclusion of a large number of optical surfaces, potentially interfering with efficient light collection, and a more limited set of options for sensor coverage.
As shown in FIG. 8, in one design of the conventional art, an array of overlapping PMTs is arranged over small 2D crystal arrays. In this case, four quadrants from four sensors cover each crystal array. However, for this design, there is no means to manage the edges and the concept seems to be better adapted to covering large surfaces. Further, modern PET systems with ToF capabilities use mostly 1 and 1½ inch PMTs, which implies a large number of very small arrays. In addition, this design has no equivalent implementation in a long and narrow detector. The 1D version shown in the top of FIG. 8, with one sensor sharing half of its surface with two adjacent sub-arrays simply does not carry enough information to perform 2D positioning.
Quadrant sharing can be implemented in two ways.
In the independent module quadrant sharing approach, adjacent independent modules, such as the module shown in FIG. 10A, are not optically coupled and do not share any PMTs. As shown in FIG. 10B, scintillator arrays of independent quadrant sharing modules cannot be placed closer than approximately the width of one photosensor (PMT). The large gaps between adjacent modules in the independent quadrant sharing approach are a major disadvantage of that approach. Moreover, the independent module quadrant sharing approach on a long and narrow array would be terribly inefficient and would clearly prevent close packing.
In the continuous quadrant sharing approach, as shown in FIG. 10C, the detector is essentially continuous around the entire circumference of the ring. In this approach, scintillator arrays of quadrant sharing modules can be placed closer if they share PMTs, thereby losing their independence. In this case, the modules essentially become a single continuous detector in which some PMTs are shared between modules. The loss of independence of the modules in the continuous quadrant sharing approach means that the detector is only fully functional once every detector module has been installed, and modules cannot be fully tested and calibrated before installation. The continuous quadrant sharing approach does not allow for the production and testing of fully functional independent units prior to inclusion of the units in the scanner, and also does not allow for replacement of an independent module when part of the detector fails. This is a significant practical disadvantage for the continuous quadrant sharing approach.